Changes in ultrasound properties of porcine kidney tissue during heating

Ultrasound in Med. & Biol., Vol. 27, No. 5, pp. 673– 682, 2001 Copyright © 2001 World Federation for Ultrasound in Medicine & Biology Printed in the USA. All rights reserved 0301-5629/01/$–see front matter

PII: S0301-5629(01)00354-4

● Original Contribution CHANGES IN ULTRASOUND PROPERTIES OF PORCINE KIDNEY TISSUE DURING HEATING A. E. WORTHINGTON* and M. D. SHERAR† *Princess Margaret Hospital/Ontario Cancer Institute, and †Departments of Medical Biophysics and Radiation Oncology, University of Toronto, Toronto, Ontario, Canada (Received 14 August 2000; in final form 22 January 2001)

Abstract—Changes in the ultrasound (US) attenuation and backscatter of fresh pig kidney were measured as the tissue was heated. The objective was to use these changes to predict how an US image would change in real-time with a view to its use as a monitoring tool for minimally invasive thermal therapy (MITT). Separate samples of fresh pig kidney were heated from 37°C to temperatures of 45°, 50°, 55°, 60° and 65° with warm water. Measurements were made over the frequency range from 3.5 MHz to 7.0 MHz during 30-min heating experiments. A general increase in attenuation magnitude (dB/cm) and slope (dB/cm-MHz) was observed at temperatures of 55°C or greater. Little change in backscatter power was observed during heating to 45°C. At higher temperatures, the changes in backscatter showed a more complex pattern throughout the experiments, but still showed a trend of increase to a greater value at the end of heating than at the start. This backscatter increase was greater at higher temperatures. The net effect of the changes in US properties suggests that it may be possible to use diagnostic US to monitor, in real-time, MITT in kidney. (E-mail: [email protected]) © 2001 World Federation for Ultrasound in Medicine & Biology. Key Words: Ultrasound, Ultrasonic tissue characterization, Backscatter, Attenuation, Thermal therapy, Cancer.

(RF) (McGovern et al. 1999), microwave (Kigure et al. 1996), laser (Williams 1999), intracavitary photon radiation sources (Chan et al.1999) and shock waves (Oosterhof et al. 1990) have been reported. Most of the studies have been conducted in small animals, with some early human experience in laparoscopic energy delivery. Minimally invasive thermal therapy (MITT) is a promising experimental treatment, in which the target tissue is heated to temperatures above 50°C to rapidly cause coagulation (Thomsen 1991; Thomsen et al. 1989; Pearce et al. 1991). With MITT, it is highly desirable to find a real-time method to monitor tissue damage to spare healthy tissue from the rapidly spreading heat. The clinician also wishes to monitor if the thermal lesion size and shape matches that of the tumor, to ensure that all the cancer cells are killed. Diagnostic ultrasound (US) could be excellent for monitoring MITT because it is noninvasive and the images are displayed in real-time. Several studies have investigated the change in US properties when tissue is heated. Gertner et al. (1997) studied changes in attenuation and backscatter of beef liver during heating to temperatures of up to 70°C. Below 55°C, normalized backscatter initially rose, then fell to below its initial value. Above 55°C, the effect was reversed.

INTRODUCTION Small masses are being discovered incidentally with increasing frequency in the kidneys of asymptomatic human patients (Mevorach et al. 1992; Wuderlich et al. 1998). For the last 50 years, the standard treatment has been radical nephrectomy (Robson et al. 1969). Recently, urologists have questioned the need to sacrifice full renal units to deal with these slow-growing masses that have not yet metastesized (Bosniak1995). Partial nephrectomy or nephron-sparing surgery (NS) is increasingly being performed as an alternative and is showing five-year survival rates of over 80% (Morgan and Zincke 1990). Laparoscopic NS has been employed, but control of bleeding continues to be a challenge (Ono et al. 1999). Minimally invasive therapies (MIT), designed to ablate small renal masses and avoid incision of the vascular renal parenchyma, are, therefore, being investigated. Studies using cryotherapy (Nakada et al. 1998), hightemperature thermal therapy using high-intensity focused ultrasound (HIFU) (Visioli et al. 1999), radiofrequency Address correspondence to: M. D. Sherar, Head Radiation Physics, 610 University Ave., Toronto, Ontario, Canada M5G 2M9. E-mail: [email protected] 673

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Normalized attenuation showed an initial drop that, above 55°C, was followed by a rise. Gertner et al. (1998) also measured a permanent drop in echogenicity as beef liver was heated to increasing temperatures. Malone et al. (1992) generated real-time images of pig liver treated with interstitial laser photocoagulation (ILP) and observed increased echogenicity around the fibreoptic tip. This effect was much more pronounced when vaporization was produced. Bush et al. (1993) reported a large increase in US backscatter in a region of coagulated necrosis when the tissue had not been degassed, and a small reduction in backscatter when the tissue had been degassed. These studies indicate that it is possible to observe changes in an US image of tissue as it is being heated. This study presents changes in the US attenuation and backscatter of fresh porcine kidney measured during heating. These parameters were measured over a frequency range of 3.5 MHz to 7.0 MHz. The intention was to use these changes to predict the appearance of realtime US images during MITT, although it should be kept in mind that this is an in vitro study, whereas MITT will be delivered in vivo. The experiments presented are similar to those of Gertner et al. (1997), with some differences. Gertner and colleagues used a fixed 3.5-MHz transducer to collect a single A-scan from a sample of bovine liver tissue. The fixed transducer arrangement required that the heating system be reassembled for each of the eight samples of tissue for which results were averaged in that work. The system presented here moved a 5.0-MHz transducer in a X-Y raster that allowed the collection of 36 uncorrelated A-scans in each sample of porcine kidney tissue. MATERIALS AND METHODS The experimental apparatus is shown in Fig. 1. The aluminum tissue holder and the Plexiglas temperature chamber were cylindrically symmetrical. The high-temperature water chamber was used to heat the tissue sample to the temperature of interest. The tissue sample itself was kept in phosphate-buffered saline (PBS) to ensure osmotic equilibrium, and was connected thermally to the high-temperature water chamber via the thin Mylar威 membrane and the aluminum reference plate. The larger water tank was kept at 37°C to minimize the volume of hot water that was necessary to heat the tissue; thus, allowing improved control over the tissue temperature. The transducer was mounted on a 2-D motion system (Parker Hannifin Corp., Harrison City, PA). The motion controller card was mounted in a personal computer (PC) and independently controlled the two axes of motion, each of which consisted of a cross roller slide and a stepping motor. The steps per revolution and pitch of the

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Fig. 1. Schematic of experimental apparatus. The large tank was kept at 37°C throughout the experiment. Initially, the high-temperature water chamber (and the tissue) was brought to 37°C. Then, the high-temperature water bath was set to ⬃80°C and hoses switched to bypass the chamber. This resulted in a reservoir of hot water which, when the hoses were switched back to the chamber at heating time 0, heated the chamber and tissue very quickly. The location of the thermocouples (T/C) are shown. The temperature-recording system and the mechanism for moving the transducer in a X-Y raster are not shown.

drive screws were such that each step represented a travel of 1.2 ␮m. The motion system was bolted to a frame that was mounted on top of the tank. The transducer (Matec, Northborough, MA) was 6 mm in diameter, f/6, 5.0-MHz, broadband PZT with a full width half maximum (FWHM) at the focus of 1.9 mm and a depth of field of 105 mm. The transducer was driven by a 6-ns impulse from a model UTA-3 Transducer Analyzer (Aerotech Laboratories, Lewistown, PA), which resulted in the transducer emitting a broad band pulse centered at its resonant frequency of 5 MHz. The echo signal traveled through an in-house designed and built 56 dB PC-controlled attenuator, through a 1.17-MHz high-pass filter and a 7.23-MHz low-pass filter, and into an analog to digital converter (ADC) (Sonix, Springfield, VA) in the same PC as the motion control card. The ADC had an analog bandwidth of 175 MHz, produced 8-bit digital output and sampled at 100 MHz (i.e., the sample interval was 0.01 ms). Temperature control The large tank was filled with degassed water that was maintained at 37°C using a pump and temperaturecontrolled water bath (Haake F3, Karlsruhe, West Ger-

Heated kidney US properties ● A. E. WORTHINGTON and M. D. SHERAR

many). The water in this tank was previously degassed by connecting a vacuum flask to an aspirator on a cold water tap. Another pump and Haake F3 were connected to the high-temperature water chamber in a closed loop. The water that was to be circulated through the hightemperature water chamber was first boiled to thoroughly degas it, and allowed to cool before being loaded. This was done to minimize any bubbles that might form in the circulating water when it was heated to high temperature. Initially, the circulating water in the high-temperature water chamber was also brought to 37°C. When all the parts of the system reached 37°C, the hoses were switched so that the circulating water bypassed the hightemperature water chamber (Fig. 1), after which the high-temperature water bath was set to ⬃80°C. This resulted in a reservoir of very hot water which, when the hoses were switched back to the high-temperature water chamber at the start of heating, heated it and the tissue to the desired final temperature very rapidly. The tissue temperature was closely monitored and kept at the final desired high temperature by adjusting the temperature setting of the high-temperature water bath. The temperatures of the tank, the high-temperature water chamber, and both temperature-controlled water baths were measured using 0.65 mm type K thermocouples encased in 1.25-mm Teflon威. Thermocouple readings were acquired by a Labmate Data Acquisition and Control System (Sciemetric, Ottawa, Ontario, Canada). The Labmate was controlled by a PC that also recorded and displayed the temperatures. Tests (presented later) confirmed that temperature readings from the thermocouple in the hightemperature water chamber agreed with readings from a thermocouple inserted through the membranes directly into the tissue. Tissue preparation Porcine kidneys were obtained directly from a slaughterhouse to ensure freshness and were brought within the hour to our lab to be used immediately or to be immediately stored in PBS at 4°C. The majority of the experiments were performed immediately or within 24 h. Two experiments were performed at 48 h. Measurements of attenuation and backscatter power at 37°C were made as a function of post mortem time to ensure the validity of the data. Each experiment required a fresh sample of tissue. Due to the use of a 37°C baseline, it was not possible to perform multiple heating experiments on the same piece of tissue. The porcine kidneys used in this study were typically 8 ⫻ 16 ⫻ 3.5 cm. The cortex was homogeneous in appearance and formed a layer of ⬃1 cm thickness around the medullary pyramids. Sections from the cortex were sliced to thicknesses of slightly greater than 5 mm. Lateral dimensions were typically 30 or 40 mm. These kidney slices were then placed in PBS

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in a vacuum flask connected to a vacuum pump that provided an absolute pressure of about 50 mmHg. for 30 min, to ensure that the tissue was degassed. The tissue holder was placed at the bottom of a small tub, into which the degassed PBS and tissue were poured. The tissue was approximately centered in the holder to be covered with a Mylar威 membrane (25 micron thick, ⬃150 mm diameter). All manipulations of the tissue and membrane were performed while they were submerged in the PBS. This was necessary to ensure that no air was trapped under the membrane or reabsorbed into the tissue. The membrane was stretched tight and clamped over the tissue by pressing an aluminum membrane clamping ring down onto the tissue holder (Fig. 1). Eight machine screws were then used to bolt the ring to the holder. The tissue holder was mounted and sealed into the hightemperature water chamber, which was then filled with the circulating water. The high-temperature water chamber was then placed in the tank, and both the tank water and the circulating water were brought up to 37°C. Data acquisition A 100 ⫻ 100 mm raster was performed over the tissue holder with A-scan data being recorded at 2-mm intervals. This interval was chosen to ensure that adjacent RF lines (A-scans) were not correlated. For each A-scan, the signal amplitude from the central 1.5 mm of tissue was squared and integrated. Values were stored and displayed as a C-scan to allow the selection of a region of interest (ROI), over which a more detailed examination was undertaken. The ROI was chosen in a region that was visually homogeneous in appearance in the C-scan. The selected ROI was a 10-mm square with US information being collected in a raster with an increment of 2 mm. This gave 36 positions in the ROI at which A-scans were collected. At each position, the PC-controlled attenuator was switched on and four Ascans were sampled, averaged and recorded. Then the attenuator was switched off and a further 128 A-scans were sampled, averaged and recorded with the transducer in the same position. A single raster scan over the ROI required approximately 23 s and was done at the start of every minute of the 45-min experiment. Data were collected at 37°C for 15 min. The start of the raster scan at 15 min was defined as time zero in the graphs of attenuation and backscatter vs. heating time presented later in this paper. Heating was started immediately upon completion of this raster scan. Data were then collected for 30 min of heating. Figure 2 shows representative A-scan data for one point in a ROI. The transducer was approximately 38 mm (50 ␮s round trip) from the tissue holder. Therefore, time zero in Fig. 2 does not correspond to the launch of the pulse from the transducer. The echo from the mem-

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PBS and the speed of sound in the tissue. These speeds were used to calculate the thickness of the tissue in the ROI. The tissue holder was machined cylindrically symmetrical, as shown in Fig. 1. The time difference between the echoes from the 25-␮m Mylar威 membrane and the tissue holder, taken at the holder edge where the physical separation of the two echoes was known, was used to calculate the speed of sound in PBS. This speed of sound was used to calculate the physical separation of the echoes through PBS close to the tissue, which, in turn, was used with the time difference through tissue to calculate the speed of sound in the tissue. A value for the average thickness of tissue over the ROI was then calculated to be used in the attenuation calculations. Attenuation The frequency-dependent attenuation coefficient ␣(␯) (dB cm ⫺1) was derived from the treatment of D’Astous and Foster (1986) and is given by: Fig. 2. Typical RF data collected at each point in the ROI. The data in the right half of the figure is from the same path through the tissue as in the left half. The left half is attenuated to show the echoes from the Mylar威 membrane and from the tissue holder. The right half is not attenuated, to show the backscatter from the tissue. The transducer is ⬃38 mm (50 ␮s round trip) in front of the tissue holder and is, therefore, off-scale to the left. The echo from the Mylar威 membrane on the front of the high-temperature water chamber is also off-scale to the left. The reverberations were monitored to ensure they did not overlap the tissue or tissue holder echoes.

brane on the front of the high-temperature water chamber is off-scale before time zero on this graph. The data in the first half of Fig. 2 is with the PC-controlled attenuator on and shows the echoes from the Mylar威 membrane at the front of the tissue and from the tissue holder. The echoes from the tissue are too small to be seen on this scale. The data in the second half is identical information starting at the same distance from the transducer, but is shown with the attenuator off to display the echoes from the tissue. Here, the echoes from the Mylar威 membrane and holder are off-scale in magnitude. Two reverberations are noted in Fig. 2. One was between the membrane on the high-temperature water chamber and the transducer and the other was between the two membranes (Fig. 1). Speed of sound While the tissue was at 37°C, and after the tissue had reached its final temperature, the attenuator was switched on and a reference set of A-scans was taken at 2-mm increments across the full width of the tissue holder through the middle of the ROI. These reference A-scans were used to calculate the speed of sound in the

␣ 共 ␯ 兲 ⫽ ␣ PBS共 ␯ 兲 ⫹

冉

冊

20 S PBS共 ␯ 兲 . log10 2D S H共 ␯ 兲

(1)

Here ␣PBS(␯) is the temperature-dependent attenuation of PBS, which was taken to be the same as in water (Duck1990), SH(␯) and SPBS(␯) are the amplitude spectra for the signal reflected by the tissue holder through tissue and PBS, respectively and D is the average thickness of the tissue in the ROI calculated according to the method described in the previous section. The amplitude spectrum through PBS was taken by averaging the fast Fourier transform (FFT) of 3 or 4 echoes from the tissue holder through PBS close to the tissue. These echoes were taken from the reference set of A-scans described in the speed of sound section. The amplitude spectrum through tissue was the average of the FFTs of the tissue holder echoes collected at each of the 36 positions in the ROI. Frequency-dependent attenuation coefficients were calculated from 3.5 MHz to 7.0 MHz in steps of 0.25 MHz. All calculations were repeated each min of each heating experiment. Backscatter The change in backscatter coefficient with heating time was calculated by measuring the ratio BSCt/BSC0 where BSCt and BSC0 are the backscatter coefficients at time t and time 0, respectively. This ratio, averaged over the 36 points of the ROI, was calculated at each frequency from 3.5 MHz to 7.0 MHz in steps of 0.25 MHz. The US backscatter coefficient was derived from the treatment of Turnbull et al. (1989) and is shown in eqn (2).

Heated kidney US properties ● A. E. WORTHINGTON and M. D. SHERAR

BSC ⫽

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R 2 ␲ 共1 ⫺ cos⌰兲

冕

␯⫹0.25MHz

兩S K共 ␯ ⬘兲 ⫻ 10 D␣共␯⬘兲/ 20兩 2d ␯ ⬘

⫻ ␯⫺0.25MHz ␯⫹0.25MHz

冕

⫻

1 . 共 z 2 ⫺ z 1兲

(2)

兩S PBS共 ␯ ⬘兲兩 2d ␯ ⬘

␯⫺0.25MHz

Here, R is the intensity reflectance of the PBS/aluminum tissue holder interface and ⌰ is the half angle subtended by the transducer at its focus. SK is the amplitude spectrum of the backscatter signal from the kidney tissue at temperature T and SPBS is the amplitude spectrum of the reference echo, through PBS, from the aluminum tissue holder at temperature T. These spectra were integrated over a range ␯ ⫺ 0.25 MHz to ␯ ⫹ 0.25 MHz in steps of 0.05 MHz. SK was multiplied by the factor 10D␣(␷)/20 shown in eqn (2) to compensate for the frequency-dependent attenuation of kidney tissue. This factor represents the attenuation through the average distance traveled through tissue by the RF signal. Because the minimum and maximum distances traveled through tissue were 0 and twice the thickness, respectively, the average distance equals the thickness D. This distance represents a compromise between over- and undercompensation for attenuation in eqn (2). It was necessary to gate the backscatter signal to include only echoes from the tissue. Matlab威 (MathWorks, Natick, MA) was employed to select the start and end of the gate by displaying expanded sections of a typical set of unattenuated RF data (the right side of Fig. 2). The start and end of the gate were chosen to include as much tissue signal as possible without including echoes from the Mylar威 membrane or from the tissue holder. The depth of tissue, z2 ⫺ z1, specified by the length of the gate in time would change as the speed of sound in kidney changes. For this reason, it was necessary to scale the gate length in time by the ratio of the speeds of sound measured before and during heating. It was not necessary to evaluate the first term in eqn (2), because all backscatter coefficients were normalized to their values at t0, where t0 is at the start of heating. Note that, because these calculations square the amplitude spectra, relative backscatter is a ratio of powers. RESULTS Thermal history The thermal history of a typical 60°C heating experiment is shown in Fig. 3. One thermocouple was inserted into the water of the high-temperature water

Fig. 3. Thermal history of the high-temperature water chamber and pig kidney during heating. One thermocouple was placed in the water inside the high-temperature water chamber and one was inserted through the Mylar威 membranes directly into the tissue. Temperatures were sampled at 5-s intervals.

chamber and another, for this test only, was inserted directly into the tissue through the Mylar威 membranes of the temperature chamber and the tissue holder. Both the chamber and tissue probes rose to 90% of the target temperature in approximately 75 s. The two probes were in agreement to within 1°C for the remainder of the experiment, indicating that the temperature measured in the chamber was a good surrogate for the tissue temperature. Speed of sound Figure 4 shows the speed of sound in pig kidney, in the PBS adjacent to the pig kidney and in clean PBS for

Fig. 4. Speed of sound in pig kidney, in PBS adjacent to the pig kidney, and in clean PBS. For the kidney and PBS adjacent to kidney, the points are the averages of 21, 4, 3, 3, 3 and 3 experiments at 37°, 45°, 50°, 55°, 60° and 65°C, respectively. For PBS only, all values are the average of 5 experiments. The error bars are SD.

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Fig. 5. Attenuation coefficient (E) in dB/cm and backscatter power (●) in arbitrary units for pig kidney at 37°C plotted vs. time post mortem. The values are the averages of 8, 8, 4 and 3 experiments at 1 h, 24 h, 48 h and 72 h, respectively. The error bars are the SEM.

temperatures from 37°C to 65°C. The values presented for tissue and for the PBS adjacent to the tissue were calculated 15 min after the start of heating for those temperatures above 37°C. Because each experiment had a 37°C component, the 37°C value was averaged over 20 experiments. The speed of sound in pig kidney was 1571 m/s at 37°C, increased to 1591 at 50°C, and decreased to 1545 m/s at 65°C. The speed of sound for the PBS adjacent to the tissue was 1545 m/s at 37°C, increased to 1585 m/s at 55°C, and then decreased to 1545 m/s at 65°C. The speed of sound in PBS alone was 1541 m/s at 37°C and rose to 1595 m/s at 65°C. Attenuation The attenuation coefficient spectra at 37°C were plotted vs. time post mortem in Fig. 5. The graphs show no significant differences as function of time post mortem. Attenuation coefficients as a function of frequency from 3.5 MHz to 7.0 MHz measured at 45°, 50°, 55°, 60° and at 65°C are shown in Fig. 6. For each experiment, the values at 36 positions in the raster scan over the ROI were averaged. The data points presented are the average of at least three experiments, as specified in the caption. The numerical data are presented in Table 1. The center frequency of the transducer was 5.0 MHz. This was approximately the mid point of the frequency range examined. The graphs of attenuation coefficient (␣) (Fig. 6) showed little change during heating at 45°C and 50°C. At 55°C ␣ (5.0 MHz) showed a small increase by a factor of 1.17 at 15 min, followed by a slight dip at 30 min while the slope steadily increased with time by a factor of 1.7.

Fig. 6. Least square fit of attenuation coefficient of pig kidney vs. frequency at (a) 45°, (b) 50°, (c) 55°, (d) 60° and (e) 65°C at several heating times. The plots are the averages of 4, 3, 3, 3 and 3 experiments at 45°, 50°, 55°, 60° and 65°C, respectively. The error bars are SEM.

At 60°, ␣ (5.0 MHz) increased by a factor of 1.36 and the slope more than doubled during the 30-min experiment. The greatest changes were observed in the 65°C experiments, where ␣ (5.0 MHz) and the attenuation slope increased by factors of 1.86 and 3.05, respectively, over 30 min. The attenuation coefficient at 5.0 MHz. was plotted vs. heating time at different temperatures from 45°C to 65°C, with each graph being normalized to its own value at 0 min (Fig. 7). Each graph represents the average of at least 3 experiments, as specified in the caption. The 45°C, 50°C and 55°C curves show slight rises to 1.2 over

Table 1. Attenuation coefficient at 5.0 MHz (dB/cm) and slope (dB/cm-MHz) for experimental temperatures and heating times Temperature 45° 50° 55° 60° 65°

0 min

2 min

15 min

30 min

2.033, 0.42 1.475, 0.38 1.584, 0.37 1.271, 0.21 1.402, 0.22

2.001, 0.38 1.662, 0.38 1.767, 0.44 1.495, 0.25 1.800, 0.31

2.066, 0.45 1.561, 0.48 1.857, 0.56 1.556, 0.46 2.151, 0.56

2.099, 0.47 1.549, 0.52 1.796, 0.64 1.732, 0.56 2.601, 0.67

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Fig. 7. Attenuation at 5.0 MHz plotted vs. heating time at several temperatures. Each curve is normalized to its value at time 0. The plots are the averages of 4, 3, 3, 3 and 3 experiments at 45°, 50°, 55°, 60° and 56°C, respectively. No error bars are shown to minimize clutter.

the 30 min of heating. The 60°C and 65°C curves show rises to 1.45 and 1.75, respectively. Backscatter Backscatter power for 5.0 MHz (in arbitrary units) at 37°C is plotted vs. time post mortem in Fig. 5. These values were not normalized so that a comparison could be made between experiments performed at different times. There was little variation from 1 h to 48 h. The normalized backscatter power at 5.0 MHz., plotted vs. heating time, is shown in Fig. 8. For each experiment, the values at 36 positions in the raster scan over the ROI were averaged. Each graph was the average of at least three experiments, as specified in the caption. At 45°C, the backscatter showed a slight increase by a factor of 1.15, followed by a slight decrease over the duration of the experiment. During experiments at temperatures of 50°C and 55°C, the backscatter rose, dipped, and then rose again by final factors of 1.55 and 1.7 respectively. At 60°C, the normalized backscatter dropped initially to 0.8, followed by a rise to 2.2 times the original value. At the highest temperature, 65°C, the normalized backscatter increased by a factor of approximately 2.5 after 10 to 15 min, after which there was a slight decrease. DISCUSSION Speed of sound Figure 4 shows the speed of sound for pig kidney increasing with temperature up to a peak of 1590 m/s at 50°C, after which it decreased. Most published data regarding US properties of tissues were measured at 40°C or lower (Duck 1990). Few authors have published

Fig. 8. Backscatter power at 5.0 MHz from pig kidney vs. heating time, normalized at start of heating. (a) 45°, (b) 50°, (c) 55°, (d) 60° and (e) 65°C. The plots were the averages of 4, 3, 3, 3 and 3 experiments at 45°, 50°, 55°, 60° and 65°C, respectively. The error bars are SEM.

results at higher temperatures (Bamber and Hill 1979). Bamber and Hill found that the speed of sound in nonfatty tissues increased up to a maximum at about 50°C. Duck (1990) compiled similar results from a number of studies in several nonfatty tissues, also suggesting that the temperature coefficient may become negative above 50°C. The values for speed of sound in the PBS adjacent to the tissue rose from 1545 m/s at 37°C to 1558 m/s at 55°C, and then decreased at higher temperatures. This decrease at higher temperatures was unexpected. Coppens (1981) presented an equation for the speed of sound in saline. The values for 0.9% saline rose steadily from 1532 m/s at 37°C to 1566 m/s at 65°C. A separate measurement in clean PBS was made to resolve this discrepancy. Figure 4 shows the speed of sound in clean PBS rising from 1541 m/s at 37°C to 1595 m/s at 65°C. These values are, at most, 2% higher than Coppen’s values, assuming 0.9% saline. During a tissue-heating experiment, particularly at higher temperatures, the PBS adjacent to the tissue became visibly contaminated. This suggests that blood, and possibly other fluids, were being flushed out of the tissue as it coagulated. Duck (1990)

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states that, for most liquids, particularly lipids (which can be a significant component of blood), velocity drops with temperature up to ⬃50°C. It is possible that this trend continues above 50°C and is the cause of the high-temperature decrease in speed shown in Fig. 4. It should be noted that the speed of sound in the PBS adjacent to the tissue was of interest in this experimental setup only as far as it was used to calculate the speed of sound in tissue and the thickness of tissue in the ROI. These values are used in the calculations of attenuation and backscatter. Attenuation Little change was observed in attenuation coefficient during heating experiments at 45°C and 50°C. A general increase in attenuation coefficient with heating time was observed at temperatures of 55°C and greater (Figs. 6, 7). This agrees with previous bovine liver (Gertner et al. 1997; Bamber and Hill 1979) and canine liver (Damianou et al. 1997) data. Gertner et al. (1997) found an initial decrease in attenuation during the first 2 min of heating, followed by an increase at later times, except at their highest temperature, 70°C. It was hypothesized that the initial drop was due to a temperature effect, after which the tissue coagulated, causing an increase in attenuation due to a structural effect. In the 70°C experiment of Gertner et al.(1998), the coagulation occurred too rapidly to allow the initial drop to be observed. The results presented in this paper do not show an initial decrease (Figs. 6, 7). The time to reach the final temperature in this paper was ⬃1.25 min compared to ⬃7 min in the experiments by Gertner and colleagues. This much quicker temperature rise may have led to the tissue coagulating much more quickly than in the previously published experiments. Here, the initial drop in attenuation due to the temperature effect may have occurred too quickly to have been observed. One other difference observed here with porcine kidney compared to bovine liver is a greater increase in slope of attenuation vs. frequency at higher temperatures (Fig. 6). Figure 5 shows insignificant changes in attenuation at 5.0 MHz for up to 72 h post mortem. This is a longer storage time than was used in this work. Crosby and Mackay (1978), in their examination of bovine spleen at 4.0 MHz, demonstrated a constant attenuation for up to 50 h post mortem, followed by a steady drop. Their tissue was kept cool and moist with ice in a closed container, but was not degassed before the experiments. They suggested that the drop at 50 h was caused by deterioration of the tissue. This deterioration would have resulted in the formation of bubbles that would have increased the absorption and backscatter. The insignificant changes in attenuation shown in Fig. 5 for times as long as 72 h post mortem are most likely because the

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pre-experiment degassing removed bubbles that had formed in the tissue. Bamber et al. (1977) made backscatter (discussed below) and attenuation measurements on several types of tissue that they stored in tap water at room temperature for up 120 h. These storage temperature and media were chosen to be the “worst possible situation.” The tissues were degassed before examination and displayed insignificant changes to attenuation over the times involved. Backscatter A general increase in relative backscatter was observed as temperature and heating time increased (Fig. 8). The standard error (SE) of the measurements also became larger as heating time and temperature increased. The tissue was always well degassed before each experiment, but it is possible that the sudden initial heating produced some gas bubbles that were then reabsorbed. This could result in the initial rise followed by a fall in the backscatter, as is seen in Fig. 8a, b and c. The final rise in backscatter (Fig. 8b and c) could be caused by the gas bubbles being permanently formed in the tissue. Tissue at 60°C and 65°C (Fig. 8d and e) is definitely coagulated and would not be expected to reabsorb any gas bubbles. It would be expected that fluid would be driven out of the tissue as it coagulates, and that the backscatter would rise. The initial dip in Fig. 8d remains unexplained. The results of Gertner et al. (1997) show some similarities, in that at 50°C and 55°C, backscatter rises then drops but, at higher temperatures, it drops then rises. Figure 5 shows insignificant changes in backscatter power for post mortem times of up to 48 h (the maximum storage time used in this paper) and suggests that values at 72 h might be unreliable. Bamber et al. (1977) show large changes in the backscatter amplitude of bovine spleen and porcine liver, which they attribute to the autolytic changes that may begin immediately after excision due to their “worst possible situation” of tissue storage and media. The tissue in this study was kept refrigerated, which should be expected to slow any deterioration and delay changes to the backscatter. Also note that this study presents backscatter power, whereas Bamber and colleagues present amplitude measurements. Ultrasound image appearance during MITT This study examined only in vitro porcine kidney where blood would be expected to coagulate along with the tissue. In the in vivo case, the flowing blood may not coagulate in larger vessels. In a living body, the kidney is encased in a tough membrane, the capsule, which prevents it from expanding due to the high volume of blood (⬃25% of the cardiac output) flowing through it. During a thermal therapy, the physiologic response is to

Heated kidney US properties ● A. E. WORTHINGTON and M. D. SHERAR

increase blood flow, which will carry more heat away from the heated region but, in the case of the kidney, its size or density will not change. However, with these caveats, it is of interest to predict image changes during MITT in the kidney based on these results. Figure 7 shows that, at a temperature insufficient to cause coagulation, there would be little change in attenuation during a 30-min exposure. At temperatures that cause coagulation, Fig. 7 demonstrates an increase in attenuation coefficient during a 30-min exposure. Figure 6d and e show that, at higher frequencies, the attenuation increase is even greater. The net effect would be to further decrease the amount of acoustic energy reaching deeper tissue. This would lead one to expect a decrease in brightness in the treated area on an US image. This decrease would be noticeable in time, and as a shadowing caused by the thermal lesion. At the same time, there would be an increase in brightness due to the increase in backscatter. Figure 8 shows that this backscatter increase is delayed in time with respect to the increase in attenuation. The net effect should be one of decreased echogenicity during an MITT coagulation of tissue, as found by Gertner et al. (1998). All the tissue used in these experiments was well degassed. If a MITT produced bubbles in the target tissue, they would shadow the treatment field, making real-time monitoring difficult (Rendon et al. 2000). Figures 7 and 8 show that changes in attenuation and backscatter may not be apparent for several min. If the MITT technique employed were one in which the heat was delivered to the target tissue over a very short time, the target tissue might cool down before any observable ultrasonic property changes could occur. An examination of ultrasonic properties of coagulated tissue that had been allowed to cool back down to 37°C might be more relevant to rapid heat delivery techniques such as HIFU. If high-power US were considered as the heat source for MITT, the time- and temperature-dependent changes in attenuation observed in this work would necessitate the inclusion of varying attenuation parameters in any pretreatment US field calculations (Kolios et al. 1999). Failure to include changing attenuation in a pretreatment calculation can result in significant errors in the predicted temperature (Kolios et al. 1999). CONCLUSIONS The intent of MITT is to increase the temperature of a targeted tissue to ⬃60°C, a temperature sufficient to cause coagulation. This paper presents time and temperature measurements of US attenuation and backscatter, which suggest that it might be possible to monitor some forms of MITT in real-time using diagnostic US using

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attenuation mapping. Real-time monitoring based on imaging of these property changes would be restricted to techniques that deliver heat to the target over several min. The ultrasonic property changes presented in this work do not occur quickly enough to allow the monitoring of very rapid delivery techniques. It must be kept in mind that this study examines in vitro samples of kidney. The effect of in vivo blood may cause significant differences in these results. Acknowledgements—This work was funded by the National Cancer Institute of Canada with funds from the Canadian Cancer Society and by Photonics Research Ontario. The authors thank Mark Gertner for help during the experiments and data processing.

REFERENCES Bamber JC, Hill CR. Ultrasonic attenuation and propagation speed in mammalian tissues as a function of temperature. Ultrasound Med Biol. 1979;5:149 –157. Bamber JC, Fry MJ, Hill CR, Dunn F. Ultrasonic attenuation and backscattering by mammalian organs as a function of time after excision. Ultrasound Med Biol 1977;3:15–20. Bosniak MA. Observation of small incidentally detected renal masses. Semin Urol Oncol 1995;13:267–272. Bush NL, Rivens I, ter Haar GR, Bamber JC. Acoustic properties of lesions generated with an ultrasound therapy system. Ultrasound Med Biol 1993;19:789 – 801. Chan D, Koniaris L, Magee C, et al. Feasibility of ablating normal renal parenchyma by intracavitary photon radiation energy in a canine model. J Endourol 1999;13:A14. Coppens AB. Simple equations for the speed of sound in Neptunian waters. J Acoust Soc Am 1981;69:862– 863. Crosby BC, Mackay RS. Some effects of time post-mortem on ultrasonic transmission through tissue under different modes of handling. IEEE Trans Biomed Eng 1978;BME-25(1):91–92. Damianou CA, Sanghvi NT, Fry FJ, Maass-Moreno R. Dependence of ultrasonic attenuation and absorption in dog soft tissues on temperature and thermal dose. J Acoust Soc Am 1997;102:628 – 634. D’Astous FT, Foster FS. Frequency dependence of ultrasound attenuation and backscatter in breast tissue. Ultrasound Med Biol 1986; 12:795– 808. Duck FA. Physical properties of tissue. London: Academic Press, 1990:75–95. Gertner MR, Wilson BC, Sherar MD. Ultrasound properties of liver tissue during heating. Ultrasound Med Biol 1997;23:1395–1403. Gertner MR, Worthington AE, Wilson BC, Sherar MD. Ultrasound imaging of thermal therapy in vitro liver. Ultrasound Med Biol 1998;24:1023–1032. Kigure T, Harada T, Yuri Y, Satoh Y, Yoshida K. Laparoscopic microwave thermotherapy on small renal cell tumors: Experimental studies using implanted VX-2 tumors in rabbits. Eur Urol 1996;30: 377–382. Kolios MC, Sherar MD, Hunt JW. Temperature dependent tissue properties and ultrasonic lesion formation. Advances in heat and mass transfer in biotechnology. ASME 1999;44:113–118. Malone DE, Wyman DR, Moote DJ, et al. Sonographic changes during hepatic interstitial laser photocoagulation. Invest Radiol 1992;27: 804 – 813. McGovern FJ, Wood BJ, Goldberg SN, Mueller PR. Radio frequency ablation of renal cell carcinoma via image guided needle electrodes. J Urol 1999;161:599 – 600. Mevorach RA, Segal AJ, Tersegno ME, Frank IN. Renal cell carcinoma: Incidental diagnosis and natural history: Review of 235 cases. Urology 1992;39:519 –522. Morgan WR, Zincke H. Progression and survival after renal conserving surgery of renal cell carcinoma: Experience in 104 patients and extended follow-up. J Urol 1990;144:852– 858.

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Nakada SY, Lee FT, Warner T, Chosy SG, Moon TD. Laparoscopic cryosurgery of the kidney in the porcine model: An acute histological study. J Endourol 1998;12(6):567–73. Ono Y, Kinikawa T, Hattori R, et al. Laparoscopic radical nephrectomy for renal cell carcinoma: A five-year experience. Urology 1999;53: 280 –286. Oosterhof GON, Smiths GEHJ, De Ruyter AE, Schalken JA, Debruyne FMJ. In vivo effects of high energy shock waves on urological tumors: An evaluation of treatment modalities. J Urol 1990;144: 785–789. Pearce JA, Cheong W-F, Pandit K, McMurray T, Thomsen S. Kinetic models for coagulative processes: Determination of rate coefficients in vivo. Lasers in dermatology and tissue welding. SPIE 1991;1422:27–33. Rendon RA, Gertner MR, Sherar M, Asch MR, Trachtenberg J, Tsilhlias J, Robinette M, Sweet J, Jewett M. Development of a radiofrequency technique for ablation of small renal masses. 55th Annual meeting of the Canadian Urology Association, Kelona, Canada, June 25–29, 2000.

Volume 27, Number 5, 2001 Robson CJ, Churchill BM, Anderson W. The results of radical nephrectomy for renal cell carcinoma. J Urol 1969;101:297–301. Thomsen S. Pathologic analysis of photothermal and photomechanical effects of laser-tissue interactions. Photochem Photobio 1991;53(6):825– 835. Thomsen S, Pearce JA, Cheong W-F. Changes in birefringence as markers of thermal damage in tissues. IEEE Trans Biomed Eng 1989;36(12):1174 –1179. Turnbull DH, Wilson SR, Hine AL, Foster FS. Ultrasonic characterization of selected renal tissues. Ultrasound Med Biol 1989;15: 241–253. Visioli AG, Rivens IH, ter Haar GR, Horwich A, Huddart RA, Moskovic E, Padhani A, Glees J. Preliminary results of a phase I dose escalation clinical trial using focused ultrasound in the treatment of localized tumours. Eur J Ultrasound 1999;9:11–18. Williams JC. Dosimetry, ultrasound and histopatholgic correlation of laser induced thermotherapy of renal cell carcinoma in man. J Endourol 1999;13:A101. Wunderlich H, Shumann S, Jantitzky V, et al. Increase of renal cell carcinoma incidence in Central Europe. J Urol 1998;33:538 –541.